Method and system for rapid magnetic resonance imaging of gases with reduced diffusion-induced signal loss

ABSTRACT

A methodology, system and computer program product for designing and optimizing a rapid magnetic resonance imaging pulse sequence for creating images of a gas or gas-filled structure with substantially reduced diffusion-induced signal attenuation during the course of data acquisition compared to that for currently available magnetic resonance imaging techniques is disclosed. The methodology and system allows desirable combinations of image signal-to-noise ration, spatial resolution and temporal resolution to be achieved that were heretofore not possible. For example, magnetic resonance imaging of hyperpolarized noble gases, which recently has shown significant promise for several medical imaging applications, particularly imaging of the human lung, can be improved. Pulse sequences designed according to the subject methods permit signal levels to be achieved that are up to ten times higher than those possible with the gradient-echo methods now commonly used for hyperpolarized-gas imaging. This signal increase can be traded for substantially lower does, and hence much lower cost, of the hyperpolarized-gas agent. The methodology and system will also be useful for non-biological applications of hyperpolarized gases for example material science studies, as well as for magnetic resonance imaging of any other gas for biological or non-biological applications. Pulse sequences designed according to the subject methods can also serve as the foundation for a variety of specialized gas-imaging pulse sequences, such as those for apparent-diffusion-coefficient, dynamic or oxygen-concentration imaging.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority from U.S. Provisional Application No.60/380,760, filed May 15, 2002, entitled “Method and Apparatus for RapidMagnetic Resonance Imaging of Gases with Reduced Diffusion-InducedSignal Loss,” the entire disclosure of which is hereby incorporated byreference herein.

FIELD OF THE INVENTION

The present invention relates to magnetic resonance imaging (MRI), andmore particularly for using hyperpolarized gases together with rapid MRIpulse sequences designed to yield reduced diffusion-induced signal loss.

BACKGROUND OF THE INVENTION

Nuclear magnetic resonance imaging (MRI) is an important modality forboth clinical and basic-science imaging applications. A recent notableadvance in MRI was the introduction of the “hyperpolarized” noble gaseshelium-3 (³He) and xenon-129 (¹²⁹Xe) as novel magnetic-resonancecontrast agents [1]. Nuclear polarization levels approaching 100 percentcan be achieved using hyperpolarized noble gases, and this dramaticincrease in the polarization compared to that typically achieved atthermal equilibrium (at most approximately 10⁻⁴) has presented theopportunity for many new MRI applications. For example, high-resolutionMR images of the lung air spaces have been demonstrated following theinhalation of hyperpolarized-³He gas [2-5], and studies suggest that ³Helung imaging shows promise for differentiating healthy lungs from thosewith pathologies such as chronic obstructive pulmonary disease [6,7],asthma [8] and cystic fibrosis [9].

Achieving a high signal-to-noise ratio (SNR) through optimization of theMRI acquisition method (that is, the “pulse sequence”) has long been afundamental goal in the development of MRI because thethermal-equilibrium nuclear magnetic resonance signal is inherentlyweak. Although hyperpolarized gases intrinsically provide a largenuclear polarization, the SNR performance of the associated pulsesequences is still of prime concern because hyperpolarized gases areexpensive to prepare and, in the case of ³He, are in limited worldsupply. Therefore, MRI pulse sequences for hyperpolarized-gas imagingthat provide high SNR of are significant practical importance.Furthermore, since the hyperpolarized magnetization is inherently in anon-equilibrium state, the lifetime of the hyperpolarized state (asmeasured by the T1 relaxation time) is limited, for many practicalapplications to 10-100 seconds, and thus applicable MRI pulse sequencesshould also acquire the image data rapidly.

In the case of sufficiently long T2 relaxation times, it is wellestablished in conventional proton (¹H) MRI that pulse sequences whichmaintain the phase coherence of at least a significant fraction of thetransverse magnetization during the application of successiveradio-frequency (RF) pulses are useful for rapid, high-SNk magneticresonance imaging. Examples of such techniques commonly used for ¹H MRIinclude RARE imaging [10] and its derivatives such as HASTE [11], andFISP imaging [12]. However, the application of these establishedtechniques for rapid, high-SNR hyperpolarized-gas imaging is limited dueto diffusion-induced signal attenuation that results from the diffusionof the gas in the magnetic-field gradients required for imaging andthose progresses during data acquisition. (The diffusivities of the freegases are approximately 10⁴-10⁵ larger than that for water protons inthe body of an animal or human.) The degree of signal attenuationincreases with decreasing voxel size (that is, increasing spatialresolution) and thus the spatial resolution is limited by the associatedimage blurring that result from the progressive signal attenuationduring data acquisition.

A recently published study that investigated the use of RARE-type pulsesequences for hyperpolarized ³He MRI of the human lung claimed, based ontheoretical analysis and corresponding experimental results, that thediffusion-dependent resolution limit for RARE-type techniques is 6 mm[13]. In contrast, transverse-magnetization-spoiled, gradient-echo-basedMRI pulse sequences currently used for hyperpolarized-gas imaging of thehuman lung typically use an in-plane resolution of approximately 3 mm,and higher spatial resolution may certainly be needed for other orfuture applications of hyperpolarized gases. Nonetheless, forhyperpolarized-gas imaging, these spoiled, gradient-echo-based MRI pulsesequences yield (for equal spatial resolution) only approximatelyone-tenth of the signal that could be provided by a RARE-type pulsesequence if the diffusion-induced signal attenuation during theRARE-type pulse sequence could be made to be negligible.

Therefore, it would clearly be of significant practical importance if itwere possible to appropriately optimize pulse sequences that maintainthe phase coherence of at least a significant fraction of the transversemagnetization during the application of successive RF pulses to minimizediffusion-induced signal attenuation and therefore permit the SNRadvantage of these techniques to be realized for hyperpolarized-gasimaging in conjunction with higher spatial resolution. This SNR increasecan be traded for substantially lower dose, and hence much lower cost,of the hyperpolarized-gas contrast agent. In addition to in-vivohyperpolarized gas imaging, such optimized techniques would potentiallyalso be useful for non-biological applications of hyperpolarized gases,for example material science studies, as well as for magnetic resonanceimaging of any other gas for biological or non-biological applications.These optimized techniques could also serve as the foundation for avariety of specialized gas-imaging pulse sequences, such as those forapparent-diffusion-coefficient [14] or dynamic [15] imaging.

BRIEF SUMMARY OF THE INVENTION

The present invention comprises the methodology, computer programproduct, and system for designing and optimizing a rapid MRI pulsesequence for creating images of a gas or gas-filled structure withsubstantially reduced diffusion-induced signal attenuation during thecourse of data acquisition compared to currently available MRItechniques. The present invention thereby allows desirable combinationsof image signal-to-noise ratio, spatial resolution and temporalresolution to be achieved that were heretofore not possible. These“diffusion-optimized” gas-imaging pulse sequences are based on thefollowing design goals, applied as appropriate for the application athand: (i) spatial-encoding magnetic-field gradient waveforms aredesigned to have a zeroth moment approximately equal to zero; (ii) asappropriate, spatial and/or spatial-spectral selection magnetic-fieldgradient waveforms are designed to have a zeroth moment approximatelyequal to zero; (iii) as appropriate, flip angles for the refocusing RFpulses are set approximately equal to 180°; (iv) as appropriate, thek-space trajectory is specifically optimized to provide a low level ofdiffusion-induced signal attenuation throughout the acquisition foreither a single-shot or multi-shot acquisition; (v) spatial-encodingmagnetic-field gradient waveforms are designed to exclude as much aspossible any periods within the waveform during which the gradientamplitude is zero; (vi) as appropriate, magnetic-field gradientwaveforms are optimized to approximately minimize diffusion-inducedsignal attenuation based on the gradient-hardware specifications; (vii)as appropriate, the data sampling period and the associatedspatial-encoding gradient waveforms are considered jointly forapproximately maximizing the SNR while maintaining a predetermined levelof image blurring, and/or a predetermined level of one or more otherimage artifacts such as a susceptibility-induced artifact; and (viii) asappropriate, the order of phase-encoding is arranged to increase theoverall signal level during the acquisition compared to that for aconventional sequential phase-encoding order. These design goals may beperformed in various orders and/or with modified procedures, systems, orstructures suitable to a given application.

An application of the present invention is magnetic resonance imaging ofhyperpolarized noble gases, which recently has shown significant promisefor several medical imaging applications, particularly imaging of thelung. Pulse sequences designed according to the methods of the presentinvention permit signal levels to be achieved that are up to ten timeshigher than those possible with the gradient-echo methods now commonlyused for hyperpolarzed-gas imaging.

The present invention will also be useful for inter alia non-biologicalapplications of hyperpolarized gases, for example material sciencestudies, as well as for magnetic resonance imaging of any other gas forbiological or non-biological applications. Pulse sequences designedaccording to the methods of this invention can also serve as thefoundation for a variety of specialized gas-imaging pulse sequences,such as those for apparent-diffusion-coefficient, dynamic oroxygen-concentration imaging.

In a first aspect, the present invention feat=res a method forgenerating a pulse sequence for operating a magnetic resonance imagingsystem for imaging a region of an object, wherein at least a portion ofthe region contains gas, including but not limited to hyperpolarizednoble gas, for at least a portion of the time required to apply saidpulse sequence, said method comprising the steps of:

-   -   (a) selecting spatial-encoding magnetic-field gradient waveforms        to be approximately fully rephased, that is to have a zeroth        moment approximately equal to zero (which includes fully        rephased, that is to have a zeroth moment precisely equal to        zero), over the time period between pairs of successive said RF        pulses in said pulse sequence;    -   (b) if desired, selecting spatial and/or spatial-spectral        selection magnetic-field gradient waveforms associated with said        RF pulses in said pulse sequence to be approximately fully        rephased;    -   (c) if it is not desired to have approximately fully-rephased        spatial or spatial-spectral selection magnetic-field gradient        waveforms of step ‘b’, and/or it is desired to have spoiling        magnetic-field gradients associated with refocusing RF pulses,        setting the flip angles for said refocusing RF pulses        approximately equal to 180°;    -   (d) if a specific type of k-space trajectory is not required,        optimizing said k-space trajectory to provide a low level of        diffusion-induced signal attenuation throughout the acquisition        for either a single-shot or multi-shot acquisition; and    -   (e) selecting said spatial-encoding magnetic-field gradient        waveforms to exclude as much as possible any time in said        spatial-encoding magnetic-field gradient waveforms during which        the gradient amplitude is zero.

In a second aspect, the present invention features a magnetic resonanceimaging system for generating a pulse sequence for operating saidmagnetic resonance imaging system for imaging a region of an object,wherein at least a portion of the region contains gas, including but notlimited to hyperpolarized noble gas, for at least a portion of the timerequired to apply said pulse sequence, the system comprising:

-   -   a main magnet system for generating a steady magnetic field in        at least a region of the object to be imaged;    -   a gradient magnet system for generating temporary magnetic field        gradients in at least a region of the object to be imaged;    -   a radio-frequency transmitter system for generating        radio-frequency pulses in at least a region of the object to be        imaged;    -   a radio-frequency receiver system for receiving magnetic        resonance signals from at least a region of the object to be        imaged;    -   a reconstruction system for reconstructing an image of the        object from the received magnetic resonance signals; and    -   a control system for generating signals controlling the gradient        magnet system, the radio-frequency transmitter system, the        radio-frequency receiver system, and the reconstruction system,        wherein the control system generates signals causing:    -   (a) spatial-encoding magnetic-field gradient waveforms to be        applied that are selected to be approximately fully rephased,        that is to have a zeroth moment approximately equal to zero        (which includes fully rephased, that is to have a zeroth moment        precisely equal to zero), over the time period between pairs of        successive RF pulses in said pulse sequence;    -   (b) if desired, spatial and/or spatial-spectral selection        magnetic-field gradient waveforms associated with RF pulses in        said pulse sequence to be applied that are selected to be        approximately fully rephased;    -   (c) if it is not desired to have approximately fully-rephased        spatial or spatial-spectral selection magnetic-field gradient        waveforms of step ‘b’, and/or it is desired to have spoiling        magnetic-field gradients associated with refocusing RF pulses,        flip angles for said refocusing RF pulses to be applied that are        set to be approximately equal to 180°;    -   (d) if a specific type of k-space trajectory is not required, a        k-space trajectory to be applied that is optimized to provide a        low level of diffusion-induced signal attenuation throughout the        acquisition for either a single-shot or multi-shot acquisition;        and    -   (e) said spatial-encoding magnetic-field gradient waveforms to        be applied that are selected to exclude as much as possible any        time in said spatial-encoding magnetic-field gradient waveforms        during which the gradient amplitude is zero.

In a third aspect, the present invention features a computer readablemedia carrying encoded program instructions for causing a programmablemagnetic resonance imaging system to perform the method discussed abovein the first aspect of the invention. Similarly, the invention featuresa computer program product comprising a computer useable medium havingcomputer program logic for enabling at least one processor in a magneticresonance imaging system to generate a pulse sequence, the computerprogram logic comprising the method discussed above in the first aspectof the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing and other objects, features and advantages of the presentinvention, as well as the invention itself, will be more fullyunderstood from the following description of preferred embodiments, whenread together with the accompanying drawings in which:

FIG. 1 shows a schematic representation of the general structure of anMRI pulse sequence that maintains the phase coherence of at least asignificant fraction of the transverse magnetization during theapplication of successive RF pulses.

FIG. 2 shows theoretical calculations of the normalized signal versusecho number for a two dimensional half-Fourier RARE-type (HASTE) pulsesequence using a standard readout gradient waveform.

FIG. 3 shows theoretical calculations of the normalized signal versusecho number for a two dimensional half-Fourier RARE-type (HASTE) pulsesequence using a diffusion-optimized readout gradient waveform.

FIG. 4 shows theoretical calculations of the normalized signal versusecho number for a two dimensional half-Fourier spin-echo-train pulsesequence using a diffusion-optimized semi-circular k-space trajectory.

FIGS. 5(A)-(C) show theoretical calculations of the b value (FIG. 5(A)),waveform duration (FIG. 5(B)), and SNR characteristics of afully-rephased (zeroth moment equal to zero) readout magnetic-fieldgradient waveform as a function of the data sampling period (FIG. 5(C)),for a spatial resolution of 4 mm; thereby illustrating that an optimumconfiguration of the waveform exists with respect to diffusion-inducedsignal attenuation and showing the signal-to-noise performance for thiswaveform when implemented in a True-FISP type pulse sequence.

FIG. 6 show the theoretical (solid line) and experimental (triangles)signal decays during 128 repetitions of a True-FISP pulse sequence usingthe fully-rephased readout magnetic-field gradient waveform analyzed inFIG. 5 with a data-sampling period of 1.3 ms.

FIGS. 7(A)-(C) show coronal ³He lung MR images obtained from a healthyvolunteer using gradient-echo (FIG. 7(A)), diffusion-optimized True-FISPwith sequential phase encoding (FIG. 7(B)), and diffusion-optimizedTrue-FISP with centric phase encoding (FIG. 7(C)).

FIG. 8 illustrates a simplified exemplary embodiment of a MRI system forpracticing the present invention. The present invention method can beapplied to various commercially available MRI systems.

DETAILED DESCRIPTION OF THE INVENTION

In the following, first presented is an exemplary embodiment of a MRsystem for practicing the MR methods of the present invention forimaging an object, moving or stationary. Second, the methods of thepresent invention are described. Finally, examples of the pulse sequencedesign are described.

An Exemplary MR-System of the Present Invention

FIG. 8 illustrates a simplified schematic of a MR system 1 forpracticing the present invention. The MR system 1 includes a main magnetsystem 2 for generating a steady magnetic field in an examinationzone(s) of the MR system. The z-direction of the coordinate systemillustrated corresponds to the direction of the steady magnetic fieldgenerated by the magnet system 2.

The MR system also includes a gradient magnet system 3 for generatingtemporary magnetic fields G_(x), G_(y) and G_(z) directed in thez-direction but having gradients in the x, y or z directions,respectively. With this magnetic gradient system, magnetic-fieldgradients can also be generated that do not have directions coincidingwith the main directions of the above coordinate system, but that can beinclined thereto, as is known in the art. Accordingly, the presentinvention is not limited to directions fixed with respect to the MRsystem. In this application, for ease of description, the directions x,y and z (and the gradients along these directions) are used for thereadout direction, the phase-encode direction and slice-selectiondirection (or second phase-encode direction for 3D imaging),respectively.

Also, while traditional commercial methods provide linear gradients inthe x, y, or z directions it is also possible not to utilize all threeof these linear gradients. For example, rather than using a linear zgradient, one skilled in the art can use a z-squared dependence or someother spatial dependence to provide desired results.

The magnet systems 2 and 3 enclose an examination zone(s) which is largeenough to accommodate a part of an object 7 to be examined, for examplea part of a human patient. A power supply means 4 feed the gradientmagnet system 3.

The MR system also includes an RF transmitter system including RFtransmitter coil 5, which generates RF pulses in the examination zoneand is connected via transmitter/receiver circuit 9 to a RF source andmodulator 6.

The RF transmitter coil 5 is arranged around the part of body 7 in theexamination zone. The MR system also comprises an RF receiver systemincluding an RF receiver coil that is connected via transmitter/receivercircuit 9 to signal amplification and demodulation unit or system 10.The receiver coil and the RF transmitter coil 5 may be one and the samecoil.

A gas supply (and/or gas regulator) 8 provides hyperpolarized noble gasto the examination zone or region of the object/subject (body, cavity,or the like). The gas supply may be an attachable supply line to theobject/subject or may be a portable gas supply such as a container,bolus delivery device, or dose bag. As would be appreciated by oneskilled in the art, there are wide variety of methods and systemsadapted for supplying hyperpolarized gas to the object or subject (orregion and examination zone). For illustrative examples of magneticresonance imaging using hyperpolarized gases the following patentapplications are hereby incorporated by reference herein in theirentirety: 1) pending and co-assigned U.S. patent application Ser. No.09/804,369 filed on Mar. 12, 2001, entitled “Diagnostic Procedures UsingDirect Injection of Gaseous Hyperpolarized 129Xe and Associated Systemsand Products,” and its corresponding International Patent ApplicationSerial No. PCTI/US01/07812 filed Mar. 12, 2001 (Publication No.:WO/01/67955 A2), 2) pending and co-assigned U.S. patent application Ser.No. 09/832,880 filed on Apr. 12, 2001, entitled “Exchange-Based NMRImaging and Spectroscopy of Hyperpolarized Xenon-129,” and 3) pendingand co-assigned International Patent Application Serial No.PCT/US02/11746 filed Apr. 12, 2002, entitled “Optimized High SpeedMagnetic Resonance Imaging Method and System Using Hyperpolarized NobleGases” (Publication Nos.: WO/02/084305 A2 & A3).

Some illustrative examples of magnetic resonance imaging usinghyperpolarized gases are provided in the following patent applicationsand patents and are hereby incorporated by reference herein in theirentirety: U.S. Pat. No. 5,545,396 to Albert et al., entitled “MagneticResonance Imaging Using Hyperpolarized Noble Gases;” U.S. Pat. No.5,785,953 to Albert et al., entitled “Magnetic Resonance Imaging UsingHyperpolarized Noble Gases;” and U.S. Pat. No. 5,789,921 to Albert etal., entitled “Magnetic Resonance Imaging Using Hyperpolarized NobleGases.”

The MR system also includes an amplification and demodulation unit orsystem 10, which, after excitation of nuclear spins in a part of thebody placed within the examination space by RF pulses, after encoding bythe magnetic-field gradients and after reception of the resulting MRsignals by the receiver coil, derives sampled phases and amplitudes fromthe received MR signals. An image reconstruction unit or system 12processes the received MR imaging signals to, inter alia, reconstruct animage by methods well-known in the art, such as by Fouriertransformation. It should be appreciated by one skilled in the art thatvarious reconstruction methods may be employed besides the FourierTransform (FT) depending on factors such as the type of signal beinganalyzed, the available processing capability, etc. For example, but notlimited thereto, the present invention may employ Short-Time FT (STFT),Discrete Cosine Transforms (DCT), or wavelet transforms (WT). By meansof an image processing unit or system 13, the reconstructed image isdisplayed, for example, on monitor 14. Further, the image reconstructionunit or system can optionally process MR navigator signals to determinethe displacement of a portion of the patient.

The MR system also includes a control unit or system 11 that generatessignals for controlling the RF transmitter and receiver systems by meansof a modulator 6, the gradient magnetic field system by means of thepower supply means 4, an image reconstruction unit or system 12 and animage processing unit or system 13. In a preferred embodiment, thecontrol unit or system 11 (and other control elements in the MR system)are implemented with programmable elements, such as one or moreprogrammable signal processors or microprocessors, communicating overbusses with supporting RAM, ROM, EPROM, EEPROM, analog signalinterfaces, control interfaces, interface to computer-readable media andso forth. These programmable elements are commanded by software orfirmware modules loaded into RAM, EPROM, EEPROM or ROM, writtenaccording to well-known methods to perform the real-time processingrequired herein, and loaded from computer-readable media (or computeruseable medium), such as magnetic disks or tapes, or optical disks, ornetwork interconnections, removable storage drives, flash memory, or soforth. The present invention may be implemented using hardware, softwareor a combination thereof and may be implemented in one or more computersystems or processing systems, such as personal digit assistants (PDAs),for various applications, e.g., remote care and portable care practices.

In an embodiment, the control unit that directs a MR system forpracticing the present invention can be implemented with dedicatedelectronic components in fixed circuit arrangements. In this case, thesededicated components are arranged to carry out the method describedabove. For example, the invention is implemented primarily in hardwareusing, for example, hardware components such as application specificintegrated circuits (ASICs). Implementation of the hardware statemachine to perform the functions described herein will be apparent topersons skilled in the relevant art(s).

In particular, the control unit commanded by its loaded software causesthe generation of MR signals by controlling the application of MR pulsesequences, which comprise RF-pulses, time delays and temporarymagnetic-field gradient pulses. These pulse sequences are generatedaccording to the methods of the present invention as subsequentlydescribed, and generally include 2D and 3D imaging pulse sequences andoptionally navigator pulse sequences for determining the displacement ofthe patient or material.

Furthermore, according to alternate embodiments of the presentinvention, the MR system also optionally includes various other units(not illustrated) from which the state of motion of the part of thepatient being imaged can be measured. These can include sensors directlyindicating the instantaneous state of motion of the part of the patientbeing imaged, such as a chest belt for directly indicating chestdisplacement during respiration, or MR-active micro-coils whose positioncan be tracked, or optical means, or ultrasound means, or so forth.These units can also include sensors indirectly indicating theinstantaneous state of motion of the part of the patient being imaged.For example, electrocardiogram and peripheral pulse sensors measure thetemporal progress of the cardiac cycle, and permit inference of theactual state of motion of the heart from knowledge of cardiacdisplacements associated with each phase of the cardiac cycle. Whenthese sensors are present to measure the state of motion, the controlunit need not generate navigator pulse sequences.

Moreover, the control unit or system 11 may also include acommunications interface 24. The communications interface 24 allowssoftware and data to be transferred between and among the control unitor system 11, reconstruction unit or system 12, image processing unit orsystem 13, and monitor 14 and external devices. Examples of thecommunications interface 24 may include a modem, a network interface(such as an Ethernet card), a communications port, a PCMCIA slot andcard, etc. Software and data transferred via communications interface 24are in the form of signals that may be electronic, electromagnetic,optical or other signals capable of being received by communicationsinterface 24. The signals are provided to communications interface 24via the communications path (i.e., channel) 26. The channel 26 carriessignals and may be implemented using wire or cable, fiber optics, aphone line, a cellular phone link, a RF link, IR link, Bluetooth, andother communications channels.

Some embodiments of the present invention may be implemented assoftware/firmware/hardware with various MR systems, and methods, as oneskilled in the art would appreciate. Other exemplary systems andmethods, but not limited thereto, are disclosed in the following U.S.patents, of which are hereby incorporated by reference in their entiretyherein: U.S. Pat. No. 6,281,681 B1 to Cline et al., entitled “MagneticResonance Imaging with Interleaved Fibonacci Spiral,” U.S. Pat. No.6,230,039 B1 to Stuber et. al., entitled “Magnetic Resonance ImagingMethod and System with Adaptively Selected Flip Angles,” U.S. Pat. No.5,749,834 to Hushek, entitled “Intersecting Multislice MRI DataAcquisition Method,” U.S. Pat. No. 5,656,776 to Kanazawa, entitled“Magnetic Resonance Imaging Apparatus,” U.S. Pat. No. 5,604,435 to Fooet al., entitled “Spiral Scanning Method for Monitoring PhysiologicalChanges,” and U.S. Pat. No. 5,485,086 to Meyer et al, entitled“Continuous Fluoroscopic MRI Using Spiral K-space Scanning.”

The Methods of the Present Invention

The present invention applies to the creation of images of a suitablestationary or moving gas, or a gas-filled structure, using nuclearmagnetic resonance (NMR) pulse sequence techniques. For our purpose, theterm “suitable” means any gas that possesses nuclear magnetic propertiesthat permit NMR signals to be derived from said gas. In particular, thisinvention applies to hyperpolarized gases (e.g., noble gases helium-3and xenon-129, but not limited thereto), wherein we define the“hyperpolarized” state as a large (relative to the thermal equilibriumpolarization for the polarizable gas in the applied static magneticfield), non-equilibrium nuclear polarization created by any method,including but not limited to optical pumping and spin exchange [16].

The present invention provides the framework for the design of a familyof rapid MRI gas-imaging pulse sequences that maintain the phasecoherence of at least a significant fraction of the transversemagnetization during the application of successive RF pulses and arespecifically optimized to reduce the degree of signal attenuationresulting from the diffusion of said gas in the magnetic-field gradientsrequired for imaging. In this manner, the present invention permitsdesirable combinations of spatial resolution, image SNR and temporalresolution to be achieved that were heretofore not possible usingexisting MRI pulse sequences. In the following, we will refer to pulsesequences that are designed according to the methods of the presentinvention as “diffusion-optimized”.

Turning now to FIG. 1, FIG. 1 shows a schematic representation of thegeneral structure of an MRI pulse sequence that maintains the phasecoherence of at least a significant fraction of the transversemagnetization during the application of successive RF pulses. This basicpulse-sequence structure provides the foundation for thediffusion-optimized pulse sequences that are subject of the presentinvention. For example, in the case of a RARE-type pulse sequence, α isthe excitation RF pulse and is, typically equal to 90°, the pulses β₁,β₂ and β₃ are the first three of n refocusing RF pulses, τ₁ typicallyequals one-half of τ₂, and τ_(i) typically equals τ_(i+1) for i greaterthan one. As another example, in the case of a FISP-type pulse sequenceusing fully-rephased magnetic-field gradients over the repetition time,α is a preparation RF pulse and is typically equal to one half of β₁,the pulses β₁, β₂ and β₃ are the first three of n excitation RF pulsesthat are all typically of the same flip angle, τ₁ typically equalsone-half of τ₂, and τ_(i) typically equals τ_(i+1) for i greater thanone. The RF pulse waveforms are drawn as rectangular for simplicity, butthey may be amplitude and/or phase modulated as appropriate for thedesired application. The spacing between RF pulses may be fixed or mayvary between pairs of pulses. The boxes on the G_(encode) axis,referenced as 42 and 43, symbolically denote the magnetic-field gradientwaveforms used for spatial encoding. The boxes on the G_(select) axis,referenced as 31, 32, 33 and 34, symbolically denote the optional use ofmagnetic-field gradients for spatial and/or spatial-spectral selection.The boxes on the G_(spoil) axis, referenced as 51, 52, 53, 54, 55 and56, symbolically denote the optional use of magnetic-field gradients forspoiling of some subset of the phase-coherent transverse magnetization.

The series of RF pulses and the associated magnetic-field gradients canbe repeated if necessary to collect the desired k-space data. The timingparameters and the number of echoes may vary between repetitions. Forsome applications the complete k-space data set is collected during asingle application of the RF pulse train. Additional RF pulses, gradientpulses and/or time delays to permit additional control over the imagecontrast may optionally precede the RF pulse labeled α for at least onerepetition of the RF pulse train.

Exemplary design methods to achieve a diffusion-optimized pulse sequenceare enumerated below in design method items 1-6, each method of which isaimed at yielding a low level of diffusion-induced attenuation of thesignals collected during the pulse sequence. Note that as explained indesign methods 1-6 below, for a specific application, the use ofspecific methods may not be compatible with the particular form of pulsesequence (e.g., RARE-type or FISP-type) that is chosen as the foundationfor the diffusion-optimized pulse sequence, or with particular desiredfeatures of the pulse sequence (e.g. a rectilinear k-space trajectory).

1. The spatial-encoding magnetic-field gradient waveforms areapproximately fully rephased, i.e. have a zeroth moment approximatelyequal to zero (which includes fully rephased, i.e. have a zeroth momentequal to zero), over the time period between pairs of successive RFpulses.

2. Unless inappropriate as outlined in method 3, any spatial orspatial-spectral selection magnetic-field gradient waveforms associatedwith particular RF pulses are approximately fully rephased, i.e. have azeroth moment approximately equal to zero.

3. For the case wherein a spin-echo-train pulse sequence, such as aRARE-type pulse sequence, serves as the foundation for thediffusion-optimized pulse sequence and wherein it is not desired to haveapproximately fully-rephased spatial or spatial-spectral selectionmagnetic-field gradient waveforms and/or it is desired to have spoilingmagnetic-field gradients associated with the refocusing RF pulses, forexample to avoid specific image artifacts, the flip angle for each ofthe refocusing RF pulses is approximately equal to 180°.

4. For the case that a specific type of k-space trajectory, for examplea rectilinear or spiral trajectory, is not required, the k-spacetrajectory resulting from the spatial-encoding magnetic-field gradientwaveforms is designed to traverse, at any time during the application ofsaid gradients and including but not limited to the portions of thetrajectory during which the associated NMR signal is sampled:

-   -   a. First, following the RF pulse β₁, a trajectory segment in k        space wherein the maximum distance from zero spatial frequency        is approximately zero.    -   b. Second, following the RF pulse β₂, a trajectory segment in k        space wherein the maximum distance from zero spatial frequency        is slightly larger than that for the trajectory segment        corresponding to RF pulse β₁.    -   c. Third, following the RF pulse β₃, a trajectory segment in k        space wherein the maximum distance from zero spatial frequency        is slightly larger than that for the trajectory segment        corresponding to RF pulse β₂, and so on for the remaining        trajectory segments that are required to sample the desired        region of k space.        The set of trajectory segments comprising the full k-space        trajectory for the measurement is designed to meet the sampling        requirements for the desired field-of-view in one, two or three        dimensions, and the desired spatial resolution in each of said        dimensions. An example of a k-space trajectory that fits this        description for two-dimensional imaging is a set of concentric        circles or semi-circles wherein the center approximately        coincides with the center of k space and wherein the first        trajectory has the smallest diameter, the second trajectory has        the next smallest diameter, and so forth until the last        trajectory has the largest diameter.

5. For the case of a k-space trajectory designed to meet therequirements of method 4, the order of collection of the trajectorysegments can be rearranged so that the maximum distance from zerospatial frequency is not monotonically increasing as the RF pulse trainprogresses, but instead so that the majority of trajectory segmentscorresponding to small maximum distances from zero spatial frequency areacquired early in the RF pulse train, the majority of trajectorysegments corresponding to medium maximum distances from zero spatialfrequency are acquired approximately midway through the RF pulse train,and the majority of trajectory segments corresponding to large maximumdistances from zero spatial frequency are acquired late in the RF pulsetrain. For example, considering the first, second and third trajectorysegments described in method 4, the order 1, 3, 2, . . . , as comparedto 1, 2, 3, . . . , would yield an acceptable result.

6. For the case wherein the RF pulse train α, β₁, β₂, β₃, . . . isapplied more than once to collect the desired k-space data and wherein ak-space trajectory that meets the requirements of methods 4 or 5 isused, the trajectory segments are interleaved across repetitions of theRF pulse train so that each RF pulse train begins by sampling relativelylow spatial frequencies, followed by slightly higher spatial frequenciesand so forth.

Further optimization of the diffusion-optimized pulse sequences may bewarranted in some cases. Enumerated below, as design methods 7-10, areadditional methods that can be optionally used to further reduce thediffusion-induced attenuation of the signals collected during the pulsesequence.

7. The spatial-encoding magnetic-field gradient waveforms are designedto exclude as much as possible any “dead” times in the waveforms, i.e.,periods within the waveform during which the gradient amplitude is zero.

8. Considering the gradient performance specifications of the particularhardware being used, the diffusion-induced signal attenuation associatedwith the spatial-encoding, spatial selection, spatial-spectral selectionand/or spoiling magnetic-field gradient waveforms is approximatelyminimized by selecting optimum values for the amplitude, slew rate,duration and functional form of each portion of said waveforms.

9. The data sampling period and the associated spatial-encoding gradientwaveforms are considered jointly to approximately maximize the SNR whilemaintaining a predetermined level of image blurring, and/or apredetermined level of one or more other image artifacts such as asusceptibility-induced artifact, by balancing the signal increase thatmay occur secondary to reduced diffusion-induced signal attenuation fromshorter spatial-encoding magnetic-field gradient waveform durationsagainst the increase in noise secondary to shorter data samplingperiods, and against the severity of other image artifacts that mayincrease with longer data sampling periods.

10. For k-space trajectories that use a conventional phase-encodingtable as part of the spatial-encoding process, the order ofphase-encoding steps is arranged so that the majority of phase-encodingsteps corresponding to low spatial frequencies along the associateddirection in k space are acquired early in the acquisition, the majorityof phase-encoding steps corresponding to medium spatial frequencies areacquired approximately midway through the acquisition, and the majorityof phase-encoding steps corresponding to high spatial frequencies areacquired late in the acquisition.

The aforementioned design methods may be performed in various ordersand/or with modified procedures, systems, or structures suitable to agiven application.

EXAMPLES

Practice of the invention will be still more fully understood from thefollowing examples, which are presented herein for illustration only andshould not be construed as limiting the invention in any way. Theresults from theoretical calculations and experimental measurements arepresented below to provide examples of how the pulse-sequence designmethods listed above serve to yield pulse sequences with reduceddiffusion-induced signal attenuation compared to conventional pulsesequences currently employed for gas imaging. The theoretical resultsare from computer-based numerical simulations of the magnetizationduring a pulse sequence and are based on the well-established Blochequation relationships and relevant extensions, and the relationship[17]: $\begin{matrix}{{b(t)} = {\int_{0}^{t}{{{k\left( t^{\prime} \right)}}^{2}\quad{\mathbb{d}t^{\prime}}}}} & \lbrack 1\rbrack\end{matrix}$for the b value, which quantifies the degree of diffusion-induced signalattenuation corresponding to a specific path through k space, where k(t)is the time-dependent position in k space as determined by the appliedgradients. The experimental results were acquired using a 1.5 Teslacommercial whole-body imager (Magnetom Vision, Siemens MedicalSolutions, Iselin, N.J.) modified by the addition of a broadband RFamplifier to permit operation at the ³He resonant frequency of 48 MHz.Hyperpolarized-gas studies used a flexible ³He-chest RF coil (IGCMedical Advances, Milwaukee, Wis.). ³He experiments were performed undera physician's Investigational New Drug application (IND 57,866) forimaging with hyperpolarized ³He using a protocol approved by ourinstitutional review board. ³He gas was polarized by collisional spinexchange with an optically pumped rubidium vapor using a commercialsystem (Model 9600 Helium Polarizer; Amersham Health, Durham, N.C.).

FIG. 2 shows the theoretically predicted effects of diffusion-inducedsignal attenuation in the healthy human lung for a half-FourierRARE-type (HASTE) pulse sequence. Only the effects of in-planespatial-encoding magnetic-field gradients are included. The apparentdiffusion coefficient (ADC) of 0.2 cm²/s corresponds to that for ³He ina healthy human lung. This pulse sequence used a standard readoutgradient waveform, that is, a monopolar waveform having constantamplitude during the data-sampling period. As is well known, thisconfiguration requires a preparatory readout gradient waveform, havingan area equal to one half that of the readout gradient waveform, duringthe time period between the excitation RF pulse and the first refocusingRF pulse. This pulse sequence configuration is the same as that used inreference [13], wherein a minimum resolution of 6 mm was postulated forimaging ³He gas in the lung with a spin-echo-train pulse sequence basedon an attenuation limit of 37% for the signal remaining at the end ofthe echo train. A 36-echo train was used in reference [13] and, as seenin FIG. 2, our calculations closely match their result. That is, at aresolution of 6 mm and echo number 36, the normalized signal in FIG. 2is approximately 0.37.

FIG. 3 shows the theoretically predicted effects of diffusion-inducedsignal attenuation in the healthy human lung for a half-FourierRARE-type (HASTE) pulse sequence wherein, in contrast to FIG. 2, thereadout gradient spatial-encoding waveform is optimized according todesign methods 1, 7 and 8. (As in FIG. 2, only the effects of in-planespatial-encoding magnetic-field gradients are included, and the ADC of0.2 cm²/s corresponds to that for ³He in a healthy human lung.) Notethat, based on the 37% attenuation limit of reference [13], the spatialresolution limit for a 36-echo train has been reduced by one-third from6 mm to 4 mm.

FIG. 4 shows the theoretically predicted effects of diffusion-inducedsignal attenuation in the healthy human lung for a half-Fourierspin-echo-train pulse sequence wherein, in contrast to FIGS. 2 and 3,the spatial-encoding magnetic-field gradient waveforms are optimizedaccording to design methods 1, 4, 7 and 8, and, based on method 4, asemi-circular k-space trajectory is used. (As in FIGS. 2 and 3, only theeffects of in-plane spatial-encoding magnetic-field gradients areincluded, and the ADC of 0.2 cm²/s corresponds to that for ³He in ahealthy human lung.) Note that, based on the 37% attenuation limit ofreference [13], the spatial resolution limit for a 36-echo train hasbeen reduced by more than two-thirds from 6 mm to less than 2 mm.Collectively, FIGS. 2-4 illustrate how the design methods of the presentinvention permit the diffusion-induced signal attenuation during thecourse of a spin-echo-train acquisition to be substantially reduced.

FIG. 5(A) demonstrates design method 8, the optimization of amagnetic-field gradient waveform based on the specifications for aparticular gradient hardware system. In this example, the configurationof a fully-rephased readout gradient waveform that yields the minimum bvalue, and therefore the minimal diffusion-induced signal attenuationassociated with this waveform, is derived for a spatial resolution of 4mm and for a gradient system having a maximum gradient amplitude of 24mT/m and a maximum gradient slew rate of 40 T/m/s. Naively, it may seemobvious that the minimum b value would coincide with the minimumduration of the readout gradient waveform. However, as shown bycomparing FIGS. 5(A) and 5(B), this is not the case—the minimum b valueoccurs for a waveform duration that is larger than the minimum duration.

FIG. 5(C) presents data for implementing design method 9 for a FISP-typepulse sequence using fully-rephased gradient waveforms (i.e., aso-called “True-FISP” pulse sequence), 80 phase-encoding steps and aflip angle of 70° for the RF pulses β_(i). FIG. 5(C) plots the SNRs forthe first and last phase-encoding steps in the RF pulse train, and plotsthe ratio of these values. One would like to achieve a high value forthe SNR of the first phase-encoding step, and for the ratio to maintainnegligible image blurring secondary to signal attenuation during the RFpulse train. However, this type of pulse sequence is extremelysusceptible to image artifacts from magnetic-field inhomogeneities;these artifacts can be reduced by decreasing the time period betweensuccessive RF pulses. In our case, this favors a reduction of thewaveform duration. For application at a relatively high magnetic fieldstrength, FIG. 5(C) suggests that the configuration corresponding to adata-sampling period of 1-1.5 ms would be a good choice.

The validity of these theoretical calculations is supported by theexperimental results presented in FIG. 6, for which the fully-rephasedreadout magnetic-field gradient waveform described in the previous twoparagraphs was implemented with a data-sampling period of 1.3 ms in aTrue-FISP pulse sequence on a 1.5 Tesla whole-body imager (describedabove). Projection images were acquired of a torso-shaped phantomcontaining two high-pressure gas cells (Amersham Health, Durham N.C.,USA). Each cell contained a mixture of O₂ and thermally polarized ³Heand had the following characteristics: T1/T2, 1180/630 ms; diffusioncoefficient, 0.26 cm²/s. Pulse sequence parameters included: TR/TE,6.14/3.07 ms; matrix, 128; field-of-view, 512 mm; flip angle, 70°; nophase encoding. FIG. 6 shows excellent agreement between thetheoretically calculated signal decay (due both to T2 relaxation anddiffusion-induced signal attenuation) and the experimentally measuredsignal intensities. Note that in this phantom roughly half of the signalloss is due to T2 decay. For in-vivo lung imaging, the T2 of ³He isestimated to be several seconds. Thus, for a typical healthy human lung,this gradient configuration should result in only 30% signal attenuationfrom diffusion for 128 RF pulses, making it viable for high SNR,high-quality imaging of the lung using hyperpolarized ³He.

For further experimental demonstration of the validity of our methods,the fully-rephased readout magnetic-field gradient waveform implementedin a True-FISP pulse sequence was used to obtain ³He lung images in fourhealthy human volunteers; the image quality and SNRs were compared tothose for images obtained using a conventional gradient-echo pulsesequence, which is the current standard in the art for imaging the lungwith hyperpolarized ³He gas. The ³He MRI studies in humans wereperformed as described above. Coronal lung images were acquired duringsuspended respiration using both True-FISP and gradient-echo pulsesequences. As discussed above, the readout gradient waveform in theTrue-FISP pulse sequence was optimized to provide a low b value (0.046s/cm²) for a maximum gradient slew rate of 40 T/m/s, a spatialresolution of 4 mm, and a data sampling period of 1.3 ms. Consideringboth the readout and slice-selection gradients, thetheoretically-predicted signal attenuation for 80 phase-encoding stepsand an apparent diffusion coefficient of 0.2 cm²/s was 30%. Otherparameters for the True-FISP pulse sequence included: TR/TE, 6.14/3.07ms; matrix, 80*128; field-of-view, 320*512 mm; flip angle, 70°; slicethickness, 15 mm; phase-encoding order, sequential or centric.Parameters for the gradient-echo pulse sequence were: TR/TE, 7/2.7 ms;matrix, 80-112*128; field-of-view, 315-368*420-512 mm; flip angle,9-10°; slice thickness, 10-15 mm. The SNR was calculated for theTrue-FISP and gradient-echo images, correcting for any differences inthe voxel volume and the net hyperpolarized magnetization inhaled by thesubject. Image artifacts were assessed by visual comparison.

The ratios of the mean SNRs for the True-FISP pulse sequences withsequential and centric phase-encoding to that for the gradient-echopulse sequence were 3.0±0.2 and 3.4±0.6 respectively. As expected, dueto signal decay from diffusion-induced attenuation, the SNR for theTrue-FISP pulse sequence with centric phase encoding was slightly higherthan that for sequential phase encoding. FIG. 7 shows a comparison ofgradient-echo and True-FISP images from the same subject, acquired withidentical spatial resolution. The overall depiction of the lung issimilar with all techniques, but the True-FISP method shows artifactualintensity variations near the diaphragmatic border attributable to thehigh sensitivity of this technique to field inhomogeneities and to thefact that the experiments were performed at a relatively high magneticfield strength of 1.5 Tesla. The depiction of subtle detail was alsoslightly better in the gradient-echo images, although this can also beattributed to susceptibility variations in the vicinity of the pulmonaryvessels and to the fact that the experiments were performed at arelatively high magnetic field strength of 1.5 Tesla.

One attractive feature of hyperpolarized-gas imaging is that, since themagnetization is created external to the main magnetic field using thelaser-based polarization system, the SNR should be approximatelyindependent of field strength as long as subject noise dominates coilnoise. Therefore, application of these optimized True-FISP pulsesequences at field strengths of a few tenths of a Tesla should suppressthe susceptibility-induced image artifacts while maintaining the SNRadvantage.

The theoretical and experimental results presented above demonstratethat diffusion-optimized pulse sequences designed according to themethods of the present invention can provide image SNR several timeshigher than that which can be achieved using current methods.

In conclusion, an advantage of the present invention methodology,system, and computer program product is that it provides a means fordesigning and optimizing a rapid magnetic resonance imaging pulsesequence for creating images of a gas or gas-filled structure withsubstantially reduced diffusion-induced signal attenuation during thecourse of data acquisition compared to that for currently availablemagnetic resonance imaging techniques.

Another advantage of the present invention is that it provides amethodology and system that allows desirable combinations of imagesignal-to-noise ratio, spatial resolution and temporal resolution to beachieved that were heretofore not possible. For example, one applicationof the present invention is magnetic resonance imaging of hyperpolarizednoble gases thereby achieving significant promise for several medicalimaging applications, particularly imaging of the human lung.

Further, another advantage of the present invention is that it providespulse sequences designed according to the subject methods to permitsignal levels to be achieved that are up to ten times higher than thosepossible with the gradient-echo methods now commonly used forhyperpolarized-gas imaging. This signal increase can be traded forsubstantially lower dose, and hence much lower cost, of thehyperpolarized-gas agent.

In addition, another advantage of the present invention methodology andsystem is that it will also be useful for non-biological applications ofhyperpolarized gases, for example material science studies, as well asfor magnetic resonance imaging of any other gas for biological ornon-biological applications.

Still yet, another advantage of the present invention is that itprovides pulse sequences designed according to the subject methods thatserve as the foundation for a variety of specialized gas-imaging pulsesequences, such as those for apparent-diffusion-coefficient, dynamic oroxygen-concentration imaging.

The present invention may be embodied in other specific forms withoutdeparting from the spirit or essential characteristics thereof. Theforegoing embodiments are therefore to be considered in all respectsillustrative rather than limiting of the invention described herein.Scope of the invention is thus indicated by the appended claims ratherthan by the foregoing description, and all changes which come within themeaning and range of equivalency of the claims are therefore intended tobe embraced herein.

REFERENCES

The following references as cited throughout this document are herebyincorporated by reference herein in their entirety:

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1. A method for generating a pulse sequence for operating a magneticresonance imaging system for imaging a region of an object, wherein atleast a portion of the region contains gas for at least a portion of thetime required to apply said pulse sequence, said method comprising thesteps of: (a) selecting spatial-encoding magnetic-field gradientwaveforms to be approximately fully rephased, that is to have a zerothmoment approximately equal to zero, over the time period between pairsof successive RF pulses in said pulse sequence; (b) if desired,selecting spatial and/or spatial-spectral selection magnetic-fieldgradient waveforms associated with RF pulses in said pulse sequence tobe approximately fully rephased; (c) if it is not desired to haveapproximately fully-rephased spatial or spatial-spectral selectionmagnetic-field gradient waveforms of step ‘b’, and/or it is desired tohave spoiling magnetic-field gradients associated with refocusing RFpulses, setting the flip angles for said refocusing RF pulsesapproximately equal to 180°; (d) if a specific type of k-spacetrajectory is not required, optimizing said k-space trajectory toprovide a low level of diffusion-induced signal attenuation throughoutthe acquisition for either a single-shot or multi-shot acquisition; and(e) selecting said spatial-encoding magnetic-field gradient waveforms toexclude as much as possible any time in said spatial-encodingmagnetic-field gradient waveforms during which the gradient amplitude iszero.
 2. The method of claim 1, further comprising: selecting themagnetic-field gradient waveforms to approximately minimizediffusion-induced signal attenuation based on gradient-hardwarespecifications.
 3. The method of claim 2, wherein the diffusion-inducedsignal attenuation associated with said spatial-encoding, said spatialselection, said spatial-spectral selection and/or said spoilingmagnetic-field gradient waveforms is approximately minimized byselecting optimum values for the amplitude, slew rate, duration andfunctional form of each portion of said magnetic-field gradientwaveforms.
 4. The method of claim 1, further comprising: consideringjointly said spatial-encoding gradient waveforms and associated datasampling period for approximately maximizing the SNR while maintaining apredetermined level of image blurring, and/or a predetermined level ofone or more other image artifacts such as a susceptibility-inducedartifact.
 5. The method of claim 4, further comprising: balancing signalincrease that may occur secondary to reduced diffusion-induced signalattenuation from shorter spatial-encoding magnetic-field gradientwaveform durations against increase in noise secondary to shorter datasampling periods, and against severity of other image artifacts that mayincrease with longer data sampling periods.
 6. The method of claim 1,further comprising: arranging the order of phase-encoding to increasesaid SNR during the acquisition compared to that for a conventionalsequential phase-encoding order.
 7. The method of claim 6, furthercomprising: ensuring that for said k-space trajectories that use aconventional phase-encoding table as part of the spatial-encodingprocess, the order of phase-encoding steps is arranged so that themajority of phase-encoding steps corresponding to low spatialfrequencies along the associated direction in said k-space are acquiredearly in the acquisition, the majority of phase-encoding stepscorresponding to medium spatial frequencies are acquired approximatelymidway through the acquisition, and the majority of phase-encoding stepscorresponding to high spatial frequencies are acquired late in theacquisition.
 8. The method of claim 1, wherein step ‘c’ is employed whena spin-echo-train pulse sequence serves as the foundation for said MRIpulse sequences.
 9. The method of claim 1, wherein step ‘d’ furthercomprises: selecting a first trajectory segment in said k-spacefollowing the associated RF pulse such that the maximum distance fromzero spatial frequency is approximately zero; and selecting subsequenttrajectory segments needed to sample the desired region of said k-spacethat follow subsequent RF pulses such that the maximum distance fromzero spatial frequency is increased incrementally for each additionaltrajectory segment.
 10. The method of claim 9, wherein the order ofcollection of said trajectory segments can be rearranged so that themaximum distance from zero spatial frequency is not monotonicallyincreasing as the RF pulse train progresses.
 11. The method of claim 10,wherein the rearrangement provides that the majority of said trajectorysegments corresponding to small maximum distances from zero spatialfrequency are acquired early in said RF pulse train, the majority ofsaid trajectory segments corresponding to medium maximum distances fromzero spatial frequency are acquired approximately midway through said RFpulse train, and the majority of trajectory segments corresponding tolarge maximum distances from zero spatial frequency are acquired late insaid RF pulse train.
 12. The method of claim 9, wherein said RF pulsetrain is applied more than once to collect the desired k-space data andfurther comprises interleaving said trajectory segments acrossrepetitions of said RF pulse trains so that each said RF pulse trainbegins by sampling relatively low spatial frequencies and then continuesby sampling increasingly higher spatial frequencies.
 13. The method ofclaim 1, wherein said gas is selected from at least one ofhyperpolarized noble gases helium-3 and xenon-129.
 14. A magneticresonance imaging system for generating a pulse sequence for operatingsaid magnetic resonance imaging system for imaging a region of anobject, wherein at least a portion of the region contains gas for atleast a portion of the time required to apply said pulse sequence, thesystem comprising: a main magnet system for generating a steady magneticfield in at least a region of the object to be imaged; a gradient magnetsystem for generating temporary magnetic field gradients in at least aregion of the object to be imaged; a radio-frequency transmitter systemfor generating radio-frequency pulses in at least a region of the objectto be imaged; a radio-frequency receiver system for receiving magneticresonance signals from at least a region of the object to be imaged; areconstruction system for reconstructing an image of the object from thereceived magnetic resonance signals; and a control system for generatingsignals controlling the gradient magnet system, the radio-frequencytransmitter system, the radio-frequency receiver system, and thereconstruction system, wherein the control system generates signalscausing: (a) spatial-encoding magnetic-field gradient waveforms to beapplied that are selected to be approximately fully rephased, that is tohave a zeroth moment approximately equal to zero, over the time periodbetween pairs of successive RF pulses in said pulse sequence; (b) ifdesired, spatial and/or spatial-spectral selection magnetic-fieldgradient waveforms associated with RF pulses in said pulse sequence tobe applied that are selected to be approximately fully rephased; (c) ifit is not desired to have approximately fully-rephased spatial orspatial-spectral selection magnetic-field gradient waveforms of step‘b’, and/or it is desired to have spoiling magnetic-field gradientsassociated with refocusing RF pulses, flip angles for said refocusing RFpulses to be applied that are set to be approximately equal to 180°; (d)if a specific type of k-space trajectory is not required, a k-spacetrajectory to be applied that is optimized to provide a low level ofdiffusion-induced signal attenuation throughout the acquisition foreither a single-shot or multi-shot acquisition; and (e) saidspatial-encoding magnetic-field gradient waveforms to be applied thatare selected to exclude as much as possible any time in saidspatial-encoding magnetic-field gradient waveforms during which thegradient amplitude is zero.
 15. A magnetic resonance imaging system forgenerating a pulse sequence for operating said magnetic resonanceimaging system for imaging a region of an object, wherein at least aportion of the region contains gas for at least a portion of the timerequired to apply said pulse sequence, the system comprising: a mainmagnet means for generating a steady magnetic field in at least a regionof the object to be imaged; a gradient magnet means for generatingtemporary magnetic field gradients in at least a region of the object tobe imaged; a radio-frequency transmitter means for generatingradio-frequency pulses in at least a region of the object to be imaged;a radio-frequency receiver means for receiving magnetic resonancesignals from at least a region of the object to be imaged; areconstruction means for reconstructing an image of the object from thereceived magnetic resonance signals; and a control means for generatingsignals controlling the gradient magnet means, the radio-frequencytransmitter means, the radio-frequency receiver means, and thereconstruction means, wherein the control system generates signalscausing: (a) spatial-encoding magnetic-field gradient waveforms to beapplied that are selected to be approximately fully rephased, that is tohave a zeroth moment approximately equal to zero, over the time periodbetween pairs of successive RF pulses in said pulse sequence; (b) ifdesired, spatial and/or spatial-spectral selection magnetic-fieldgradient waveforms associated with RF pulses in said pulse sequence tobe applied that are selected to be approximately fully rephased; (c) ifit is not desired to have approximately fully-rephased spatial orspatial-spectral selection magnetic-field gradient waveforms of step‘b’, and/or it is desired to have spoiling magnetic-field gradientsassociated with refocusing RF pulses, flip angles for said refocusing RFpulses to be applied that are set to be approximately equal to 180°; (d)if a specific type of k-space trajectory is not required, a k-spacetrajectory to be applied that is optimized to provide a low level ofdiffusion-induced signal attenuation throughout the acquisition foreither a single-shot or multi-shot acquisition; and (e) saidspatial-encoding magnetic-field gradient waveforms to be applied thatare selected to exclude as much as possible any time in saidspatial-encoding magnetic-field gradient waveforms during which thegradient amplitude is zero.
 16. A computer readable media carryingencoded program instructions for causing a programmable magneticresonance imaging system to perform the method of claim
 1. 17. Acomputer program product comprising a computer useable medium havingcomputer program logic for enabling at least one processor in a magneticresonance imaging system to generate a pulse sequence, said computerprogram logic comprising: (a) selecting spatial-encoding magnetic-fieldgradient waveforms to be approximately fully rephased, that is to have azeroth moment approximately equal to zero, over the time period betweenpairs of successive RF pulses in said pulse sequence; (b) if desired,selecting spatial and/or spatial-spectral selection magnetic-fieldgradient waveforms associated with RF pulses in said pulse sequence tobe approximately fully rephased; (c) if it is not desired to haveapproximately fully-rephased spatial or spatial-spectral selectionmagnetic-field gradient waveforms of step ‘b’, and/or it is desired tohave spoiling magnetic-field gradients associated with refocusing RFpulses, setting the flip angles for said refocusing RF pulsesapproximately equal to 180°; (d) if a specific type of k-spacetrajectory is not required, optimizing said k-space trajectory toprovide a low level of diffusion-induced signal attenuation throughoutthe acquisition for either a single-shot or multi-shot acquisition; and(e) selecting said spatial-encoding magnetic-field gradient waveforms toexclude as much as possible any time in said spatial-encodingmagnetic-field gradient waveforms during which the gradient amplitude iszero.